Most modern MRI systems use a superconducting solenoid to establish a uniform B.sub.O (or B.sub.Z) over the imaging volume. This results in the magnetic field being collinear with the path available for sample access. Coils are then required to produce monotonic, (preferably linear) gradients in B.sub.Z with respect to x, y, and z over the sample region during precisely determined pulse sequences. The transverse gradients (.delta.B.sub.Z /.delta.x, .delta.B.sub.Z /.delta.y) in the prior art have generally been established by symmetrically located sets of saddle coils, similar to those first described by Golay in U.S. Pat. No. 3,569,823 or by related planar coils as disclosed by Roemer, U.S. Pat. No. 4,926,125 and Morich et al, U.S. Pat. No. 5,036,282. Maxwell pairs or related geometries are universally used to generate the axial gradient. A co-pending application, Ser. No. 07/912,149, discloses the use of coil geometries more complex than surface currents to achieve order-of-magnitude improvements in several critical parameters for transverse gradient coils: acoustic noise and DC gradient efficiency.
The instant application discloses (a) the combined use of Golay-type and crescent-coil geometries for greatly improved switching efficiency and (b) the convenience of internal water cooling with transverse gradients. The closest prior art to the instant invention, in terms of magnetic field configuration, appear to be the trapezium loops for use with an electromagnet, as disclosed in the article "Magnetic Field Gradient Coils for NMR Imaging" by Bangert and Mansfield in Journal Physics, E, 15, 235 (1982), some screening concepts disclosed by Mansfield in U.S. Pat. No. 4,978,920, and the above referenced co-pending patent application.
The gradient pulses induce eddy currents and vibrations in nearby conducting structures (especially in flimsy shields, in the cryostat, and in lightweight rf coils) which perturb the field homogeneity following the pulses with time and spatial dependencies that are not easily characterized. Actively shielded coils for MRI were first publicly disclosed by Mansfield in February 1986 at approximately the same time that Roemer filed the patent application which resulted in U.S. Pat. No. 4,737,716. Prior independent work was underway at Doty Scientific, which shipped the first such commercially available NMR gradient coils in January 1987. Actively shielded dipolar coils for energy storage were previously disclosed by Westendorp, U.S. Pat. No. 3,671,902. Actively shielded, constant-gradient, quadrupolar magnetic field coils based on cylindrical current sheets for atomic beam confinement and focusing were previously disclosed by Beth, U.S. Pat. No. 3,466,499.
FIG. 1 approximately depicts the fingerprint coils of Schenck, Hussain, and Edelstein, U.S. Pat. No. 4,646,024, as used to generate .delta.B.sub.Z /.delta.y in an imaging region in the sample. Such a pattern achieves both higher linearity and higher switching efficiency than first-order Golay coils A similar set of concentric coils rotated 90.degree., is used to generate .delta.B.sub.z /.delta.x.
The major gradient design problems center on the following: (1) limited available space because of economic considerations, (2) motion-induced artifacts arising from the finite stiffness and mass of the coil support structure, (3) practicable coil winding (or etching) techniques, (4) acoustic noise abatement, (5) heat dissipation, and (6) minimization of transverse field components.
The conflicting technical requirements may be partially addressed by means of local planar gradient coils with highly nonlinear response, as disclosed by Roemer, U.S. Pat. No. 4,926,125. By adding distortion correction algorithms to the image processing, it is possible to use gradients with .+-.40% to .+-.60% non-linearity on one axis in applications where high spatial resolution is required only over a small portion of the image.
The following parameters generally need to be specified for gradient coil systems: gradient coefficient .alpha. (T/Am) (sometimes called gradient efficiency in the prior art), imaging sphere diameter d.sub.i (m) for a specified linearity deviation, inductance L (H), resistance R.sub.E (.OMEGA.), maximum continuous power dissipation P (W), maximum pulse current I.sub.P (A) in a specified B.sub.O, recovery time T.sub.D (s) for a specified pulse, acoustic noise for a specified pulse sequence in a specified field, and ratio of transverse field energy in the sample region to axial field energy in the imaging region.
For the fastest imaging technique, Echo Planar Imaging (EPI) and related techniques, the most important parameters are recovery time, gradient switching efficiency, transverse fields, and acoustic noise. Although EPI was first described more than 15 years ago, it has seldom been used because prior art gradient coils (a) may require megawatts of gradient driver power on the frequency-encoding axis, (b) generate sound pressure levels that are painful and damaging to the patient's hearing, (c) produce motion-related artifacts that cannot be fully removed even with the most sophisticated image postprocessing, and (d) require high power audio amplifiers costing up to several million dollars. A recent experimental demonstration at 0.5 T required nearly half a megawatt (at 10% duty cycle) at 1 kHz, and others have proposed the use of 2 MW at 5 kHz, 1.5 T, and 50% duty cycle for slice-interleaved EPI techniques. The above problems may be partially addressed using a tuned transverse gradient with sinusoidal (monochromatic) current; but the conventional gradient coil has very low electrical Q; and there are penalties in SNR and heat dissipation. Also, computational analysis becomes more complex, but the software is available.
While the Maxwell z-gradient is considerably more efficient than the Golay transverse gradient, the frequency-encoding gradient must be in the plane of the image, which often must be transverse for medical reasons. Therefore, improvements are needed in transverse gradients.
The image artifact problem can begin to be appreciated by noting that while the frequency-encoding gradient may be driven with a 500 kVA trapezoidal waveform, the phase-encoding gradient is being driven with short "blips" of several kilowatts at very low duty cycle, and the slice-selection axis is nulled. It is quite easy for nonlinear, vibration-dependent couplings between the frequency-encoding axis and the other axes to destroy the required degree of orthogonality between the axes and produce phase-related artifacts.
It should be pointed out that there is ambiguity in the definition and usage of the term "linearity" in the MRI gradient literature. Henceforth, we use this term to indicate the rms deviation of local field slope compared to the mean field slope over a specified volume. This definition is generally more demanding and a better indicator of image quality than the more common definition where linearity is defined as the maximum gradient field deviation relative to a linear field at any point in the sample, as the latter definition averages local fluctuations along the gradient axis. Other definitions can be less demanding and less useful.
The availability of better image processing and distortion correction techniques suggests that the rms gradient deviation a be increased to 14% compared to the more typical 10% value for many prior art gradients, to increase imaging volume. It is still important that the field be monotonic, but the method of Schenck et al in U.S. Pat. No. 4,646,024 results in relatively poor switching efficiency, intolerable acoustic noise, and unmanageable motion-related artifacts.
The enormous bandwidth (several MHz) of high-resolution EPI (and other more advanced techniques) can reduce the imaging time by two or three orders of magnitude without placing unrealistic demands on modern computers since computational power per cost has increased at the rate of 40% per year for the past seven years and that rate is expected to continue for several more years. Designing for strong gradients with larger gradient non-linearity with very fast switching places increased (though inconsequential) computational demands on the image processing. While there may be some increased variation in SNR over the final image, this is more than offset by the increased data rate.
In practice, using conventional shielded gradient coils, the inductive energy (I.sup.2 L/2) is larger than suggested by simple energy estimates by a substantial factor. In a co-pending patent application, methods are disclosed that allow increases in .alpha..sup.2 /L by factors of 2 to 10 compared to prior art Golay coils. However, for large systems the most important efficiency figure-of-merit often is .eta..sub.s =.alpha..sup.2 d.sub.i.sup.4 h.sub.i /.mu..sub.O L where h.sub.i is the imaging region field-of-view in the axial direction. The instant invention allows increases in .eta..sub.s by factors of 2 to 20 compared to prior art.
Some prior gradient coil designs have also suffered under the false notion that there is an inherent advantage with very low inductance coils. Higher inductance (more turns) requires higher voltage, but not higher power (VA) for the same switching time. In fact, reducing inductance below 100 .mu.H is detrimental as lead inductance and transmission line problems then becomes significant. Coil orthogonality (for isolation) and net force cancellation both dictate that integral number of turns be used in all coil sets and coil subsets. Hence, the accuracy of the shielding is limited from this quantization. The more turns, the more precisely the gradients can be shielded. Optimum number of turns is thus determined largely by the VA characteristics and economics of available power devices, magnetic shielding accuracy requirements, and standard wire insulation practice, making 250 V to 800 V (peak differential voltages for a balanced line) at 20 A to 300 A best for large systems. Optimum inductance is typically 0.2 to 1 mH.